Semiconductor radiation detector, positron emission tomography apparatus, semiconductor radiation detection apparatus, detector unit and nuclear medicine diagnostic apparatus

ABSTRACT

Each semiconductor radiation detector used for a nuclear medicine diagnostic apparatus (PET apparatus) is constructed with an anode electrode A facing a cathode electrode C sandwiching a CdTe semiconductor member S which generates charge through interaction with γ-rays. Then, a thickness t of the semiconductor member S sandwiched between these mutually facing anode electrode A and cathode electrode C is set to 0.2 to 2 mm. Furthermore, the devices are mounted (laid out) on substrates in such a way that the distance (distance of conductor) between the semiconductor radiation detector and an analog ASIC which processes the signal detected by this detector is shortened. Furthermore, the substrates on which the detectors are mounted are housed in a housing as a unit (detector unit).

CROSS-REFERENCE TO RELATED APPLICATION

The present application is related to a U.S. Serial No. ______ beingfiled based on Japanese Patent Application No. 2003-342437 filed on Sep.30, 2003, the entire content of which is incorporated herein byreference.

BACKGROUND OF THE INVENTION

The present invention relates to a nuclear medicine diagnosticapparatus, and more particularly, to a positron emission tomography(hereinafter referred to as “PET”) apparatus, which is a kind of anuclear medicine diagnostic apparatus using a semiconductor radiationdetector, semiconductor radiation detection apparatus or detector unit.

A detector using a NaI scintillator is known as a conventional radiationdetector for detecting radiation such as γ-rays. With a gamma camera (akind of nuclear medicine diagnostic apparatus) provided with a NaIscintillator, radiation (γ-rays) incident on the scintillator at anangle restricted by many collimators interacts with NaI crystals andemits scintillation light. This light travels in such a way as tosandwich a light guide, reaches a photoelectric multiplier and becomesan electrical signal. The electrical signal is shaped by a measuringcircuit mounted on a measuring circuit fixing board and transferred froman output connector to an external data collection system. All thesescintillator, light guide, photoelectric multiplier and measuringcircuit, measuring circuit fixing board, etc., are housed in a lightshielding case and shielded from electromagnetic waves other thanexternal radiation.

Since a gamma camera using a scintillator has a structure with a largephotoelectric multiplier (also called “photomultiplier”) placed afterone large crystal such as NaI, its position resolution remains on theorder of 10 mm. Furthermore, since the scintillator detects radiation inmulti-stages of conversion from radiation to visible light, from visiblelight to electrons, it has a problem of having considerably poor energyresolution. For example, there is a PET apparatus (positron emissiontomography apparatus) having position resolution of 5 to 6 mm or ahigh-end PET apparatus having position resolution of 4 mm or so, butsince their photoelectric multipliers use vacuum tubes, it is difficultto further improve position resolution.

There are radiation detectors for detecting radiation according toprinciples different from those of such a scintillator, such assemiconductor radiation detectors provided with a semiconductorradiation detection element using a semiconductor material such as CdTe(cadmium telluride), TlBr (thallium bromide) and GaAs (galliumarsenide).

This semiconductor radiation detector is attracting attention becauseits semiconductor radiation detection element converts electrical chargeproduced by interaction between radiation and the semiconductor materialto an electrical signal, and therefore it has better efficiency ofconversion to an electrical signal than the scintillator and can also beminiaturized.

-   -   [Patent Document 1] JP-A-2003-79614 (paragraph No. 0016)    -   [Patent Document 2] JP-A-2003-167058 (paragraph No. 0020, 0023)

Meanwhile, when a semiconductor material such as Tl making up asemiconductor radiation detection element interacts with radiation in asemiconductor radiation detector, holes having positive electricalcharge and electrons having negative electrical charge are generated.While mobility of electrons is relatively large, mobility of holes isrelatively small. That is, electrons move relatively easily and holesmove with difficulty. This takes more time for holes to reach anelectrode than electrons. Moreover, holes may be annihilated beforereaching the electrode. This involves a problem that the detectionsensitivity of radiation is worsened. Thus, these problems requiresolutions.

It is an object of the present invention to provide a semiconductorradiation detector capable of improving detection sensitivity.

SUMMARY OF THE INVENTION

In order to solve the above described problems, a first embodiment ofthe present invention improves detection sensitivity by shortening adistance between electrodes for charge collection of a semiconductorradiation detector. That is, the distance between an anode electrode andcathode electrode or the thickness of a semiconductor area sandwichedbetween the anode electrode and cathode electrode is 0.2 to 2 mm. Inthis structure, the distance from positions of electrons and holesgenerated by interaction between the semiconductor material andradiation to the electrodes is shortened, and therefore the timerequired for them to reach the electrodes is shortened. Furthermore,shortening the distance up to the electrodes reduces the probabilitythat holes may be annihilated midway the distance.

A second embodiment of the present invention is a nuclear medicinediagnostic apparatus comprising a plurality of unit substrates includinga plurality of semiconductor radiation detectors for introducingradiation and an integrated circuit for processing radiation detectionsignals output from the plurality of semiconductor radiation detectors.This allows the semiconductor radiation detectors and the integratedcircuits which process the outputs to be disposed close to one another,with the result that when weak output signals of the semiconductorradiation detectors are transmitted to the integrated circuits, it ispossible to reduce influences of noise on the weak output signals.

The semiconductor radiation detector, analog LSI (Large Scale IntegratedCircuit), AD converter and digital LSI are preferably arranged on theunit substrate in that order and the respective elements are connectedby wiring so that a signal detected by the semiconductor radiationdetector is processed by the analog LSI, the signal processed by theanalog LSI is processed by the AD converter, and the signal processed bythe AD converter is processed by the digital LSI. By shortening thedistance between the semiconductor radiation detector and analog LSI inparticular, this structure can shorten the wiring distance between thesemiconductor radiation detector and analog LSI and thereby reduce noisesuperimposed on the wiring until the signal detected by thesemiconductor radiation detector reaches the analog LSI. In anembodiment which will be described later, the LSI (integrated circuit)corresponds to an ASIC. Also, the semiconductor radiation detectionapparatus corresponds to a combined substrate (detector substrate+ASICsubstrate) in the embodiment which will be described later.

According to the second embodiment, detection signals when thesemiconductor radiation detectors detect radiation are processed by anapplication-specific IC called “ASIC (Application Specific IntegratedCircuit)” and this embodiment is intended to solve an additional problemdiscovered by the inventor et al. that since the detection signalsoutput from the semiconductor radiation detectors are weak, the ASIC iseasily affected by noise. A reduction of the noise leads to substantialimprovement of detection sensitivity (count, peak value, time detectionaccuracy) by the semiconductor radiation detectors.

Different substrates are preferably used as the substrate for mountingthe semiconductor radiation detectors and the substrate for mounting theLSI. During ordinary operation, the two substrates are used incombination as a combined substrate (unified substrate) so that in theevent of trouble, only the troubled substrate can be replaced to therebyfacilitate maintenance and examination, etc.

A third embodiment of the present invention adopts a unit-typeconstruction in which a plurality of unit substrates includingsemiconductor radiation detectors and an integrated circuit are mountedin a frame in a detachable/attachable manner. Since it is only necessaryto mount a detector unit including a plurality of unit substrates on anuclear medicine diagnostic apparatus, a plurality of semiconductorradiation detectors can be mounted on the nuclear medicine diagnosticapparatus at a time. In this way, the time required to mount thesemiconductor radiation detectors on the nuclear medicine diagnosticapparatus can be shortened drastically.

The embodiment is preferably adapted so that these unit substrates canbe removed from the detector unit one by one or the whole detector unitcan be removed from the nuclear medicine diagnostic apparatus, or morespecifically, from the camera, which facilitates maintenance andexamination.

Note that many semiconductor radiation detectors are used for a nuclearmedicine diagnostic apparatus (radiological diagnostic apparatus) suchas PET, SPECT and gamma camera. For example, a PET uses a hundredthousand to several hundreds of thousands of (channels) semiconductorradiation detectors and there is a demand for shortening the timerequired to mount these many semiconductor radiation detectors on thenuclear medicine diagnostic apparatus. A fourth embodiment of thepresent invention is implemented to meet such a demand. There is also ademand for facilitating maintenance and examination of semiconductorradiation detectors.

The present invention can prevent or reduce deterioration of thedetection sensitivity of radiation using semiconductor radiationdetectors. The present invention can also prevent or reducedeterioration of signals detected by the semiconductor radiationdetectors. This allows, for example, a nuclear medicine diagnosticapparatus to obtain clear images.

Other objects, features and advantages of the invention will becomeapparent from the following description of the embodiments of theinvention taken in conjunction with the accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a perspective view showing a structure of a PET apparatus as anuclear medicine diagnostic apparatus according to this embodiment;

FIG. 2 schematically shows a cross section in a circumferentialdirection of the camera of the PET apparatus in FIG. 1;

FIG. 3 schematically shows a structure of a semiconductor radiationdetector in a minimum construction;

FIG. 4 is a graph comparing a “time-peak value curve” between a casewhere a thickness t of a semiconductor material of the semiconductorradiation detector is large and a case where it is small;

FIG. 5 is a graph schematically showing a relationship between thethickness t of a semiconductor material and a peak value (maximum value)of a semiconductor radiation detector;

FIG. 6 schematically illustrates a construction of a semiconductorradiation detector having a laminated structure of semiconductormaterials and electrodes (anodes, cathodes);

FIG. 7A is a front view of a combined substrate which combines adetector substrate and an ASIC substrate of the semiconductor radiationdetectors according to this embodiment, FIG. 7B is a side view of FIG.7A and FIG. 7C is a perspective view schematically showing aconstruction of the semiconductor radiation detector mounted on thedetector substrate in FIG. 7A;

FIG. 8 is a block diagram schematically showing an analog detectioncircuit;

FIG. 9 is a block diagram showing a schematic construction of a digitalASIC and a connection relationship between an analog ASIC and thedigital ASIC;

FIG. 10 is a perspective view quoted to illustrate a construction of adetector unit housing a plurality of semiconductor radiation detectors;

FIG. 11 is a side view of the detector unit in FIG. 10 without the sideplate;

FIG. 12A is a partially exploded perspective view of a camera when thedetector unit is mounted on the camera and FIG. 12B is a cross-sectionalview of the central part of the camera;

FIG. 13 is a perspective view showing a construction of a SPECTapparatus as a nuclear medicine diagnostic apparatus according toanother embodiment;

FIG. 14 is a block diagram schematically showing a circuit constructionof an analog ASIC of the SPECT apparatus in FIG. 13; and

FIG. 15 is a block diagram showing a schematic construction of a digitalASIC in the SPECT apparatus in FIG. 13 and a connection relationshipbetween the analog ASIC and digital ASIC.

DESCRIPTION OF THE EMBODIMENTS Embodiment 1

A nuclear medicine diagnostic apparatus which is a preferred embodimentof the present invention will be explained with reference to attacheddrawings in detail below as appropriate. The following are explanationsof the nuclear medicine diagnostic apparatus according to thisembodiment, distance between electrodes of a semiconductor radiationdetector, arrangement (layout) of elements such as analog ASIC on asubstrate, and elements applicable to this embodiment for constructionof substrate units, etc. Note that an analog ASIC refers to an ASIC(Application Specific Integrated Circuit) which is anapplication-specific IC for processing analog signals and is a kind ofLSI.

<<Nuclear Medicine Diagnostic Apparatus>>

First, the nuclear medicine diagnostic apparatus (radiologicaldiagnostic apparatus) according to this embodiment will be explained. Asshown in FIG. 1, a PET apparatus 1 as the nuclear medicine diagnosticapparatus is constructed by including a camera (image pickup apparatus)11, a data processing apparatus 12, a display apparatus 13, etc. Anexaminee is laid on a bed 14 to be photographed using the camera 11. Thecamera 11 incorporates many semiconductor radiation detectors 21 (seeFIG. 3, FIGS. 7A-7C, FIG. 10) to detect γ-rays emitted from the body ofthe examinee using semiconductor radiation detectors (hereinafter simplyreferred to as “detectors”) 21. The camera 11 is provided with anintegrated circuit (ASIC) for measuring peak values, detection times ofγ-rays and is designed to measure peak values and detection times ofdetected radiations (γ-rays). The data processing apparatus 12 includesa storage apparatus, a simultaneous measuring apparatus 12A (see FIG. 2)and a tomographic information creation apparatus 12B (see FIG. 2). Thedata processing apparatus 12 takes in data of peak values, detectiontimes of detected γ-rays and packet data including detector (channel)IDs. The simultaneous measuring apparatus 12A carries out simultaneousmeasurements based on this packet data, especially data of detectiontimes and detector IDs, identifies detection positions of 511 KeV γ-raysand stores them in the storage apparatus. The tomographic informationcreation apparatus 12B creates a functional image based on theidentified positions and displays it on the display apparatus 13.

As shown in FIG. 2, inside the camera 11, many detector units 2 (seeFIG. 10 for details) housing a plurality of combined substrates 20 (seeFIG. 7 for details) provided with many detectors 21 for detecting γ-raysemitted from the examinee are disposed circumferentially. The examineeis laid on the bed 14 and positioned at the center of the camera 11. Atthis time, the detectors are disposed so as to surround the bed 14. Thedetector unit 2 is designed to output for each detector 21 included inthe detector unit 2, peak value information of γ-rays obtained based ona detection signal when a detector 21 interact with γ-rays, timeinformation on γ-ray detection and address information (detector ID) ofeach detector 21. The constructions of the detector 21, combinedsubstrate 20 and detector unit 2 will be explained in detail later. Theexaminee is administered radiopharmaceuticals, for example,fluorodeoxyglucose (FDG) containing ¹⁸F whose half-life is 110 minutes.β-rays (annihilated γ-rays) are emitted from the body of the examineewhen positrons emitted from the FDG annihilate.

Hereafter, the characteristic parts of this embodiment will beexplained.

<<Semiconductor Radiation Detector; Distance Between Electrodes>>

First, the detector 21 applied to this embodiment will be explained. Asshown in FIG. 3, the detector 21 is constructed of a semiconductorradiation detection element (hereinafter referred to as “detectionelement”) 211 made of a tabular semiconductor material S, both sides ofwhich are covered with thin-plate (film) electrodes (anode A, cathode C)(minimum construction). Of these components, the semiconductor materialS is made up of a single crystal of any one of the above described CdTe(cadmium telluride), TlBr (thallium bromide), GaAs (gallium arsenide),etc. Furthermore, the electrodes (anode A, cathode C) are made of anyone material of Pt (platinum), Au (gold), In (indium), etc. In thefollowing explanations, suppose the semiconductor material S is obtainedby slicing a CdTe single crystal. Furthermore, suppose radiation to bedetected is 511 KeV γ-rays used for the PET apparatus.

An overview of the principle of γ-ray detection using the detector 21will be explained using FIG. 3. When γ-rays are introduced into thedetector 21 and interaction occurs between γ-rays and the semiconductormaterial S constituting the detector 21, an amount of hole and electronpairs schematically shown in the figure with “+” and “−” correspondingto the energy of γ-rays is generated. Here, a voltage (e.g., 300 V) forcharge collection is applied between the electrodes of the anode A andcathode C of the detector 21. Because of this, holes are moved attractedto the cathode C and electrons are moved attracted to the anode A. Whenholes and electrons are compared, as described in “Disclosure of theinvention” the ease of movement (mobility) of electrons is relativelylarge and therefore electrons reach the anode in a shorter time. On theother hand, the mobility of holes is relatively small and thereforeholes take more time to reach the cathode. Note that holes may beannihilated before reaching the electrode.

As shown in FIG. 4 which shows a comparison in the “time-peak valuecurve” between a case where the semiconductor material S (detectionelement 211) of the detector 21 is thick and a case where it is thin,the semiconductor material S having a smaller thickness t has a quickerrise of peak value and a higher maximum value of the peak value. Havinga quicker rise of the peak value contributes to improvement of theaccuracy of simultaneous measurement of the PET, for example.Furthermore, having a higher peak value contributes to increasing energyresolution. Thus, a smaller thickness t speeds up rising of the peakvalue and increases the peak value (the efficiency of charge collectionimproves) because the time for electrons and holes to reach theelectrodes (anode A, cathode C) (time of charge collection) isshortened. This is also because holes which would be conventionallyannihilated midway can reach the electrode (cathode C) withoutannihilation because of the shorter distance. Note that the thickness tcan also be expressed by the distance between the electrodes, anode Aand cathode C facing each other.

The thickness (distance between electrodes) t of the detection element211 is preferably 0.2 mm to 2 mm. This is because a thickness t of notless than 2 mm slows down the rising speed of the peak value and reducesthe maximum value of the peak value as well. On the other hand, athickness t of smaller than 0.2 mm relatively increases the thickness(volume) of the electrodes (anode, cathode) and when installed on asubstrate, the proportion of the very semiconductor material S thatinteracts with radiation decreases. That is, reducing the thickness t ofthe semiconductor material S relatively increases the thickness of theelectrode which does not interact with γ-rays on one hand, and theproportion of the semiconductor material S which interacts with γ-raysrelatively decreases on the other, with the result that the sensitivityof detecting γ-rays decreases (γ-rays pass by). Furthermore, a smallerthickness t may cause more leakage current preventing a high voltagefrom being applied for charge collection.

For the same reason, the thickness t of the semiconductor material S ispreferably 0.5 mm to 1.5 mm and such a thickness t allows more reliabledetection of γ-rays and more correct measurement of the peak value, etc.

In the case of the PET apparatus 1, since it carries out simultaneousmeasurement, one of problems to be solved is to correctly measure aγ-ray detection time. For example, in FIG. 3, there is a difference in adetection time when positions at which γ-rays interact with thesemiconductor material S are closer to the cathode C and when thosepositions are closer to the anode A. That is, since the moving speed ofholes is lower, the detection time when the interaction occurs closer tothe anode A is relatively late, while the detection time when theinteraction occurs closer to the cathode C is relatively early(approximates to a real time). That is, also when γ-rays interact withthe semiconductor material S in the same detection element 211, there isa problem that the detection time changes depending on the position atwhich the interaction takes place. More specifically, when the thicknesst is large, the difference in the detection time depending on theposition at which the interaction takes place increases. Such an eventconstitutes no big problem in other fields, but it constitutes a bigproblem in the case of the PET apparatus 1, which carries outsimultaneous measurement (simultaneous counting) on the order of nsec(nanoseconds). Therefore, in this sense, too, it is possible todetermine the detection time appropriately within the above describedthickness range. The detection time is determined by the PET accordingto an LET system or CFD system.

As shown in FIG. 5 which schematically shows a relationship between thethickness t of the semiconductor material S and peak value (maximumvalue) of the detector 21, the maximum value of the peak value decreasesas the thickness t of the semiconductor material S increases. One ofreasons that the peak value decreases is that holes are annihilatedbefore reaching the electrode. When the thickness t becomes 2 mm, thepeak value of detected radiation falls short of a threshold whereby itis possible to discriminate 511 KeV γ-rays, and therefore it is notpreferable to increase the thickness t of the semiconductor material Smore than 2 mm as described above.

As shown in FIG. 6, the detector 21 includes the semiconductor materialsS (detection elements 211) laminated in five layers each sandwichedbetween the cathode C and anode A. Each layer of the semiconductormaterial S is a single layer detector 21 having the aforementionedthickness t (0.2 to 2 mm (more preferably 0.5 to 1.5 mm)). The thicknessof the anode A and cathode C is approximately 20 microns. In thedetector 21 having a laminated structure shown in this FIG. 6, thedifferent anodes A or different cathodes C are connected to a commonwire, and therefore each layer is not designed to detect radiationindependently of other layers. In other words, when γ-rays interact withthe semiconductor material S, it is not possible to discriminate whetherthe interaction takes place on the top layer or bottom layer. Of course,it is also possible to adopt a structure in which radiation is detectedby each layer. This five-layer structure is adopted because reducing thethickness t of the semiconductor material S is preferable in increasingthe rising speed of the peak value and increasing the maximum value ofthe peak value, but a small thickness t causes more γ-rays to pass by,and therefore reducing the amount of γ-rays that pass by whileincreasing the efficiency of charge collection increases interactionbetween the semiconductor material S and γ-rays (to increase a countvalue).

Adopting the detector 21 having such a laminated structure can obtain abetter peak value rising speed and an accurate peak value and increasethe number of γ-rays (count value)(increase the sensitivity) thatinteract with the semiconductor material S.

An area s of the electrode (anode A, cathode C) is preferably 4 to 120mm². An increase of the area s increases the capacity (straycapacitance) of the detector 21 and this increase in the straycapacitance makes noise easier to superimpose, and therefore the area sof the electrode is preferably as small as possible. Furthermore, chargeproduced when γ-rays are detected is partially accumulated in the straycapacitance, and therefore there is a problem that when the straycapacitance increases, the amount of charge stored in a charge amplifier24 b of an analog ASIC 24 or further an output voltage (peak value)decreases. When CdTe is used for the detector 21, its dielectricconstant is 11 and if the area s of the detector 21 is 120 mm²,thickness t is 1 mm, then the capacity is 12 pF, which is not negligibleconsidering the fact that the stray capacitance of connectors, etc. ofthe circuit is several pF. Therefore, the area s of the electrode ispreferably 120 mm² or less.

Furthermore, the lower limit of the area s of the electrode isdetermined by position resolution of the PET apparatus. The positionresolution of the PET apparatus is determined by not only the size(array pitch) of the detector 21 but also the positron range, etc., butsince the range of positron of 18F is 2 mm, setting the size of thedetector 21 to 2 mm or less is meaningless. The method of mounting sothat the area of the electrode becomes a minimum is a case where thesurface of the electrode is placed perpendicular to the radius directionof the camera 11 and from the above described consideration, the lowerlimit of one side of the electrode is 2 mm and the lower limit of thearea s of the electrode is 4 mm².

In the above described explanations, CdTe is used as the semiconductormaterial S which interacts with γ-rays, but it goes without saying thatthe semiconductor material S may also be TlBr or GaAs, etc. Moreover,the terms “laminated structure”, “upper layer” and “lower layer” havebeen used, but these terms are based on FIG. 6 and when the viewingdirection is turned by 90° toward the horizontal direction, thelaminated structure may be read as a parallel structure and top/bottommay be read as right/left, for example. Moreover, the direction ofincident γ-rays may also be upward, downward, rightward and leftward inFIG. 6. In other words, the detector 21 has a structure in which aplurality of (e.g., five) semiconductor materials S are arranged inparallel in such a way as to sandwich cathodes C and anodes Aalternately.

<<Combined Substrate; Detector Substrate and ASIC Substrate>>

A detailed structure of the combined substrate (unit substrate) 20installed in the detector unit 2 (FIG. 10) will be explained using FIGS.7A-7C. The combined substrate (semiconductor radiation detectionapparatus) 20 comprises a detector substrate (first substrate) 20A inwhich a plurality of detectors 21 are arranged and an ASIC substrate(second substrate) 20B in which a capacitor 22, a resistor 23, analogASICs 24, analog/digital converters (hereinafter referred to as “ADC”)25 and a digital ASIC 26 are arranged.

(Detector Substrate)

With reference to FIGS. 7A-7C, the detector substrate 20A provided withthe detectors 21 will be explained. As shown in FIG. 7A, the detectorsubstrate 20A has a grid-like arrangement (mounting) of a plurality ofdetectors 21 on one side of a substrate body 20 a (4 rows of 16detectors 21=horizontal 16×vertical 4=total 64 detectors). In the radiusdirection of the camera 11, the detectors 21 are arranged in four rowson the substrate body 20 a. The 16 detectors 21 in the horizontaldirection are arranged in the axial direction of the camera 11, that is,in the longitudinal direction of the bed 14. Furthermore, as shown inFIG. 7B, since the semiconductor radiation detectors 21 are arranged onboth sides of the detector substrate 20A, a total of 128 detectors 21are arranged on one detector substrate 20A. Here, as the number ofdetectors 21 to be installed increases, it is easier to detect γ-raysand it is possible to increase position accuracy when γ-rays aredetected. For this reason, the detectors 21 are disposed on the detectorsubstrate 20A as densely as possible. In FIG. 7A, when γ-rays emittedfrom the examinee on the bed 14 travel from bottom to top in the figure(direction indicated by an arrow 32, that is, radius direction of thecamera 11), arranging the detectors 21 in the left-to-right directiondensely on the detector substrate 20A is preferable because in this way,the number of γ-rays that pass by (the number of γ-rays that passthrough the gap between the detectors 21) is reduced. This increases thedetection efficiency of γ-rays and increases spatial resolution of animage obtained.

As shown in FIG. 7B, the detector substrate 20A of this embodimentarranges the detectors 21 on both sides of the substrate body 20 a, andtherefore the substrate body 20 a can be shared by both sides comparedto the case where the detectors 21 are arranged only on one side. Forthis reason, it is possible to reduce the number of substrate bodies 20a by half and arrange the detectors 21 more densely in thecircumferential direction. Moreover, as described above, the number ofdetector substrates 20A (combined substrates 20) can be reduced by half,and therefore there is a merit of saving time and trouble to mount thecombined substrates 20 in the housing 30 (see FIG. 10) which will bedescribed later.

In the above described explanations, the 16 horizontal detectors 21 arearranged in the axial direction of the camera 11, but the arrangement isnot limited to this. For example, the 16 horizontal detectors 21 mayalso be arranged in the circumferential direction of the camera 11.

As shown in FIG. 7C, each detector 21 has a laminated structure ofsingle crystals of the aforementioned thin-film semiconductor materialsS (detection elements 211). The structure and function thereof havealready been explained with reference to FIG. 6, but supplementaryexplanations will be given here. As described above, the detector 21 isprovided with the anode A and cathode C and a potential difference(voltage) of, for example, 300 V is applied between the anode A andcathode C for charge collection. This voltage is supplied from the ASICsubstrate 20B to the detector substrate 20A via the connector C1 (FIG.7A). Furthermore, the signal detected by each detector 21 is supplied tothe ASIC substrate 20B via the connector C1. Thus, on-board wiring (forcharge collection and signal exchange) (not shown) for connecting theconnector C1 and each detector 21 is provided in the substrate body 20 aof the detector substrate 20A. This on-board wiring has a multi-layeredstructure. In this embodiment, the detection elements 211 of thedetector 21 are arranged in parallel to the substrate body 20 a.However, the detectors 21 may also be provided so that the respectivedetection elements 211 are disposed perpendicular to the substrate body20 a.

(ASIC Substrate)

Then, the ASIC substrate 20B incorporating the ASIC will be explained.As shown in FIG. 7A, the ASIC substrate 20B is provided with two analogASICs 24 and one digital ASIC 26 on one side of the substrate body 20 b.Furthermore, as shown in FIG. 7B, since the analog ASICs 24 are providedon both sides of the substrate body 20 b, one ASIC substrate 20Bincludes a total of four analog ASICs 24. Furthermore, the ASICsubstrate 20B includes eight (=4×2) ADCs 25 on one side of the substratebody 20 b and sixteen ADCs 25 on both sides. Furthermore, as manycapacitors 22 and resistors 23 as the detectors 21 are arranged on bothsides of one substrate body 20 b. Furthermore, to electrically connecttheses capacitors 22, resistors 23, analog ASICs 24, ADCs 25 and digitalASIC 26, the ASIC substrate 20B (substrate body 20 b) is provided withon-board wiring (not shown) as with the above described detectorsubstrate 20A. This on-board wiring also has a laminated structure.

These elements 22, 23, 24, 25 and 26 are arranged (on-board wiring) sothat a signal supplied from the detector substrate 20A is supplied tothe capacitor 22, resistor 23, analog ASIC 24, ADC 25 and digital ASIC26 in that order.

The ASIC substrate 20B includes a connector (spiral contact) C1 which isconnected to the on-board wiring which is connected to each capacitor 22to make electrical connections to the detector substrate 20A and asubstrate connector C2 which makes electrical connections to the dataprocessing apparatus (the unit combination FPGA which will be describedlater). Note that the above described detector substrate 20A alsoincludes the connector C1 connected to the on-board wiring which isconnected to each detector 21.

(Connection structure between detector substrate and ASIC substrate)

The connection structure between the detector substrate 20A and ASICsubstrate 20B will be explained.

The detector substrate 20A and ASIC substrate 20B are connected not withtheir respective end faces (ends) facing each other but by providing anoverlap area where both ends overlap with each other and connecting theconnectors C1 in this overlap area as shown in FIG. 7B. This connectionis made in a detachable/attachable manner using screws for clamping.These connections are made for the following reason. That is, when thecombined substrate 20 made up of the detector substrate 20A and ASICsubstrate 20B connected (combined) together is supported on one end(cantilever support) or on both ends in the horizontal direction, aforce which flexes or bends the combined substrate 20 downward isapplied to the central area (connection area) of the combined substrate20. Here, in the case where both ends are the connection area wheretheir respective end faces (ends) face each other, the connection areais easily flexed or bent, which is not preferable.

With consideration given to this aspect, this embodiment connects thedetector substrate 20A and ASIC substrate 20B not with the respectiveend faces facing each other but by providing the overlap area so thatthe areas close to the ends overlap with each other as described above.This improves toughness against flexure or bending compared to theconnection with the end faces facing each other, which is preferable.Moreover, improving toughness against flexure or bending of the combinedsubstrate suppresses dislocation of the detectors 21 and preventsdeterioration of accuracy of identifying positions at which γ-rays aregenerated. As shown in FIG. 2, the camera 11 of the PET apparatus 1 isprovided with many detector units 2 (FIG. 10) including the combinedsubstrate 20 shown in FIGS. 7A-7C in a doughnut shape and these combinedsubstrates 20 disposed at positions of 3 o'clock and 9 o'clock in thehorizontal direction in FIG. 2 are liable to flexure or bending. Thus,the toughness of the combined substrates 20 against flexure or bendingbecomes important.

The detector substrate 20A and ASIC substrate 20B are electricallyconnected using the aforementioned overlap area. For this purpose, aconnector C1 (FIG. 7A) which electrically connects the on-board wiringof both the substrates 20A and 20B is provided in the respective overlapareas of the detector substrate 20A and ASIC substrate 20B shown in FIG.7B. For the connector C1, for example, a spiral contact (R) is used toimprove electrical connections. The spiral contact (R) is made of aball-shaped connection terminal contacting a spiral contactor over awide area and provides a characteristic of realizing optimal electricalconnections. Note that when the ball-shaped connection terminal isprovided on the ASIC substrate 20B side, the spiral contactor isprovided on the detector substrate 20A side, and when the ball-shapedconnection terminal is provided on the detector substrate 20A side, thespiral contactor is provided on the ASIC substrate 20B side.

Using such an electrical connection structure between the detectorsubstrate 20A and ASIC substrate 20B allows signals to be sent from thedetector substrate 20A to the ASIC substrate 20B with low loss. Notethat when loss is small, the energy resolution on the part of thedetectors 21 improves.

Furthermore, as described above, the detector substrate 20A and ASICsubstrate 20B are connected in a freely detachable/attachable manner bymeans of screws, etc. Thus, even when trouble occurs in thesemiconductor radiation detectors 21 or ASICs 24, 26, all that should bedone is just to replace the part with trouble. Therefore, thiseliminates waste that the entire combined substrate 20 must be replaceddue to trouble in that part. Moreover, since electrical connectionbetween the detector substrate 20A and ASIC substrate 20B is made by theconnector C1 such as the aforementioned spiral contactor (R), connectionor disconnection (combination or dissociation) between the substratescan be done easily.

In the above described construction, one detector substrate 20A isconnected to the ASIC substrate 20B, but it is also possible to dividethe detector substrate into a plurality of portions. For example, twodetector substrates may be connected to the ASIC substrate, eachconsisting of eight horizontal by four vertical detectors 21. Accordingto this construction, if one detector 21 has trouble, of the twodetector substrates, only the one including the faulty detector 21 needsto be replaced and it is therefore possible to reduce waste inmaintenance (cost reduction).

(Element Layout)

Then, the layout of elements such as the detectors 21, ASICs 24, 26 ofthe combined substrate 20 will be explained with reference to FIGS.7A-7C and FIG. 8.

As shown in FIG. 8, the detector 21 is connected to the analog ASIC 24through the connector C1, capacitor 22 and resistor 23 by means ofelectrical wiring (not shown) and a detection signal of γ-rays detectedby the detector 21 is passed through the capacitor 22 and resistor 23 bymeans of the electrical wiring and processed by the analog ASIC 24.Furthermore, the signal processed by the analog ASIC 24 is alsoprocessed by the ADC 25 and digital ASIC 26.

Here, the shorter the length of the circuit and length (distance) of thewiring, the better, because there is less influence of noise or lessattenuation of a signal. Furthermore, when simultaneous measurementprocessing is carried out by the PET apparatus 1, a shorter circuit orwiring is preferred because its time delay is smaller (preferablebecause the accuracy of detection time is not lost). For this reason, inorder of the detector 21, capacitor 22, resistor 23, analog ASIC 24, ADC25 and digital ASIC 26 from the center axis of the camera 11 outward inthe radius direction of the camera 11, that is, the elements 21, 22, 23,24, 25 and 26 are arranged (layout) in this embodiment as shown in FIG.7A. This order is the same as the signal processing order by theelements 21, 22, 23, 24, 25 and 26 (see FIG. 8, FIG. 9). That is, fromthe center axis of the camera 11 outward, the “detector, analogintegrated circuit, AD converter and digital integrated circuit” arearranged on the substrate in that order and wired in the same order.Thus, it is possible to transmit a weak signal detected by the detector21 to the analog ASIC 24 with the wiring length (distance) shortened.

Since the signal of the analog ASIC 24 is subjected to processing suchas amplification, it is less susceptible to influences of noise even ifthe length of wiring from the analog ASIC 24 onward is long. That is,considering noise, there is no problem even if the wiring length fromthe analog ASIC 24 onward is long. However, with lengthy wiring, thereis a delay in signal transmission and the accuracy of the abovedescribed detection time may deteriorate.

In this embodiment, since not only the detector 21 but also the analogASIC 24 and digital ASIC 26 are included in one combined substrate 20,the detector 21, analog ASIC 24 and digital ASIC 26 can be arranged inthe longitudinal direction of the bed 14, that is, the directionperpendicular to the body axis of the examinee subject to anexamination, and therefore this eliminates the need to extend the lengthof the camera (image pickup apparatus) 11 in the longitudinal directionof the bed more than necessary. It is also possible to consider thepossibility of arranging the analog ASIC 24 and digital ASIC 26 outsidein the radius direction of the annularly arranged detector group, and inthe longitudinal direction of the bed 14, but this causes the length ofthe camera 11 in the longitudinal direction of the bed to become longerthan necessary. Furthermore, semiconductor radiation detectors are usedas the detectors 21, and analog ASIC 24 and digital ASIC 26 are used assignal processing apparatuses, the length of the combined substrate 20in the longitudinal direction is shortened and the length of the camera11 in the orthogonal direction can be shortened significantly comparedto the case where a scintillator is used. Furthermore, the combinedsubstrate 20 is provided with the detector 21, analog ASIC 24 anddigital ASIC 26 in that order in the longitudinal direction thereof, andtherefore it is possible to shorten the length of the wiring connectingthem and simplify the wiring on the substrate.

Here, in this embodiment, one analog ASIC 24 is connected to 32detectors 21 to process signals obtained from the detectors 21. As shownin FIG. 8 and FIG. 9, one analog ASIC 24 is provided with 32 sets of ananalog signal processing circuit (analog signal processing apparatus) 33made up of a slow system and fast system. One analog signal processingcircuit 33 is provided for each detector 21 and connected to onedetector 21. Here, the fast system is provided with a timing pick offcircuit 24 a to output a timing signal for identifying a detection timeof γ-rays. On the other hand, the slow system is provided with apolarity amplifier (linear amplifier) 24 c, a band pass filter (waveformshaping apparatus) 24 d and a peak hold circuit (peak value holdingapparatus) 24 e connected in this order for the purpose of calculating apeak value of the detected γ-rays. Note that the slow system is named“slow” because it takes a certain degree of processing time to calculatea peak value. Reference numeral 24 b denotes a charge amplifier(preamplifier). A γ-ray detection signal output from the detector 21 andpassed through the capacitor 22 and resistor 23 is amplified at thecharge amplifier 24 b and polarity amplifier 24 c. The amplifiedgamma-ray detection signal is passed through the band pass filter 24 dand input to the peak hold circuit 24 e. The peak hold circuit 24 eholds a maximum value of the detection signal, that is, the peak valueof a γ-ray detection signal proportional to energy of the detectedγ-rays. One analog ASIC 24 is an LSI which integrates 32 sets of analogsignal processing circuits 33.

The capacitor 22 and resistor 23 can also be provided inside the analogASIC 24, but this embodiment arranges the capacitor 22 and resistor 23outside the analog ASIC 24 for reasons such as obtaining an appropriatecapacitance and appropriate resistance and reducing the size of theanalog ASIC 24. Note that the capacitor 22 and resistor 23 arepreferably disposed outside because in this way variations in theindividual capacitance and resistance are reduced.

In the analog ASIC 24 shown in FIG. 8, the output of the slow system ofthis analog ASIC 24 in this embodiment is designed to be supplied to anADC (analog/digital converter) 25. Moreover, the output of the fastsystem of the analog ASIC 24 is designed to be supplied to the digitalASIC 26.

The analog ASIC 24 and each ADC 25 are connected via one wire whichsends slow system signals corresponding to 8 channels all together.Furthermore, each analog ASIC 24 and digital ASIC 26 are connected via32 wires which send 32-channel fast system signals one by one. That is,one digital ASIC 26 is connected to four analog ASICs 24 via a total of128 wires.

The output signal of the slow system output from the analog ASIC 24 isan analog peak value (maximum value of the graph shown in FIG. 4).Furthermore, the output signal of the fast system output from the analogASIC 24 to the digital ASIC is a timing signal showing the timingcorresponding to the detection time. Of these signals, the peak valuewhich is the slow system output is input to the ADC 25 via the wire(wire uniting 8 channels into one) connecting the analog ASIC 24 and ADC25 and converted to a digital signal by the ADC 25. The ADC 25 convertsa peak value to, for example, an 8-bit (0 to 255) digital peak value(e.g., 511 KeV→255). On the other hand, a timing signal which is a slowsystem output is supplied to the digital ASIC 26 via the above describedwire connecting the analog ASIC 24 and digital ASIC 26.

The ADC 25 sends the digitized 8-bit peak value information to thedigital ASIC 26. For this purpose, each ADC 25 and digital ASIC 26 areconnected via a wire. For example, since there are sixteen ADCs 25 onboth sides, the digital ASIC 26 is connected to the ADC 25 via a totalof sixteen wires. One ADC 25 processes signals corresponding to 8channels (signals corresponding to eight detection elements). The ADC 25is connected to the digital ASIC 26 via one wire for transmission of anADC control signal and one wire for transmission of peak valueinformation.

As shown in FIG. 9, the digital ASIC 26 comprises a plurality of packetdata generation apparatuses 34 including eight time decision circuits(time information generation apparatuses) 35 and one ADC control circuit(ADC control apparatus) 36, and a data transfer circuit (datatransmission apparatus) 37, and integrates all these elements into oneLSI. All the digital ASICs 26 provided for the PET apparatus 1 receive a500 MHz clock signal from a clock generation apparatus (crystaloscillator) (not shown) and operates synchronously. The clock signalinput to each digital ASIC 26 is input to the respective time decisioncircuits 35 in all the packet data generation apparatuses 34. One timedecision circuit 35 is provided for each detector 21 and receives atiming signal from the timing pick off circuit 24 a of the correspondinganalog signal processing circuit 33. The time decision circuit 35determines the detection time of γ-rays based on the clock signal whenthe timing signal is input. Since the timing signal is based on the fastsystem signal of the analog ASIC 24, a time close to a real detectiontime can be set as the detection time (time information). The ADCcontrol circuit 36 receives a timing signal for the timing at whichγ-rays are detected from the time decision circuit 35 and identifies thedetector ID. That is, the ADC control circuit 36 stores a detector IDcorresponding to each time decision circuit 35 connected to the ADCcontrol circuit 36 and can identify, when time information is input froma certain time decision circuit 35, the detector ID corresponding to thetime decision circuit 35. This is possible because one time decisioncircuit 35 is provided for each detector 21. Moreover, after inputtingthe time information, the ADC control circuit 36 outputs an ADC controlsignal including detector ID information to the ADC 25. The ADC 25outputs the peak value information output from the peak hold circuit 24e of the analog signal processing circuit 33 corresponding to thedetector ID by converting it to a digital signal. This peak valueinformation is input to the ADC control circuit 36. The ADC controlcircuit 36 adds the peak value information to the time information anddetector ID to create packet data. The ADC control circuit 36 has thefunctions of the ADC control apparatus for controlling the ADC 25 andthe information combination apparatus for combining the detector IDinformation (detector position information), time information and peakvalue information. The information combination apparatus outputscombination information (packet information) which is digitalinformation including those three kinds of information. The packet data(including detector ID, time information and peak value information)output from the ADC control circuit 36 of each packet data generationapparatus 34 is input to the data transfer circuit 38.

The data transfer circuit 38 sends packet data which is digitalinformation output from the ADC control circuit 36 of each packet datageneration apparatus 34 to the integrated circuit (unit combination FPGA(Field Programmable Gate array)) 31 for unit combination provided forthe housing 30 of the detector unit 2 (FIG. 10, FIG. 11) which housestwelve combined substrates 20, for example, periodically. The unitcombination FPGA (hereinafter referred to as “FPGA”) 31 sends thedigital information to the data processing apparatus 12 through aninformation transmission wire connected to the connector 38.

Since the ADC 25 converts the peak value information output from thepeak hold circuit 24 e corresponding to the detector ID informationincluded in a control signal output from the ADC control circuit 36 to adigital signal, one ADC 25 is provided for a plurality of analog signalprocessing circuits 33 in one analog ASIC 24. Therefore, there is noneed to provide one ADC 25 for each of a plurality of analog signalprocessing circuits 33 and it is possible to thereby significantlysimplify the circuit construction of the ASIC substrate 20B. Also oneinformation combination apparatus for generating combination informationis enough for a plurality of analog signal processing circuits 33 in oneanalog ASIC 24, which can simplify the circuit construction of thedigital ASIC 26. Moreover, only one ADC control apparatus foridentifying detector IDs needs to be provided for a plurality of analogsignal processing circuits 33 in one analog ASIC 24, simplifying thecircuit construction of the digital ASIC 26.

In this way, packet data output from the digital ASIC 26 and includingdetector IDs for uniquely identifying (1) peak value information, (2)determined time information and (3) detector 21, one by one is sent tothe next data processing apparatus 12 (see FIG. 1) through aninformation transmission wire. The simultaneous measuring apparatus 12Aof the data processing apparatus 12 carries out simultaneous measuringprocessing (when two γ-rays with predetermined energy are detected witha time window with a set time, this processing regards these γ-rays as apair of γ-rays generated by annihilation of one positron) based on thepacket data sent from the digital ASIC 26, counts the simultaneouslymeasured pair of γ-rays as one γ-ray and identifies the positions of thetwo detectors 21 which have detected the pair of γ-rays using thosedetector IDs. When there are three or more γ-rays detection signalsdetected within the above described time window (when there are three ormore detected detectors 21 which have detected γ-rays) the dataprocessing apparatus 12 identifies the two detectors 21 into whichγ-rays are introduced first out of three or more detectors 21 using peakvalue information, etc., on those γ-ray detection signals. Theidentified one pair of detectors 21 are simultaneously measured and onecount value is generated. Furthermore, the tomographic informationcreation apparatus 12B of the data processing apparatus 12 createstomographic information on the examinee at the position whereradiopharmaceuticals are concentrated, that is, position of malignanttumor, using count values obtained by simultaneous measurement andposition information on the detectors 21. This tomographic informationis displayed on the display apparatus 13. Information such as the abovedescribed digital information, count value obtained by simultaneousmeasurement and position information on the detectors 21 and tomographicinformation are stored in the storage apparatus of the data processingapparatus 12.

According to the above described explanations, the detector substrate20A includes the detectors 21 and the ASIC substrate 20B includes thecapacitor 22, resistor 23, analog ASIC 24, ADC 25 and digital ASIC 26.However, the detector substrate (first substrate) 20A may include thedetector 21, capacitor 22, resistor 23 and analog ASIC 24, etc., and theASIC substrate (second substrate) 20B may include the ADC 25 and digitalASIC 26, etc. By the detector substrate 20A including the detectors 21and analog ASIC 24, the distance (wire length) between the detector 21and analog ASIC 24 can be further shortened. Thus, it is possible tofurther reduce influences of noise.

Furthermore, the combined substrate 20 may include three substrates(detector substrate 20A, analog ASIC substrate and digital ASICsubstrate) and they may be connected in a detachable/attachable mannerthrough their respective connectors.

In this case, of the three substrates, the detector substrate 20Aincludes the detectors 21, the analog ASIC substrate includes thecapacitor 22, resistor 23 and analog ASIC 24 and the digital ASICsubstrate includes the ADC 25 and digital ASIC 26. This structureseparates the substrate incorporating the analog circuit from thesubstrate incorporating the digital circuit to thereby prevent noise onthe digital circuit side from entering the analog circuit. Furthermore,this structure separates the substrate incorporating the analog ASICfrom the substrate incorporating the digital ASIC and connects the twosubstrates using a detachable/attachable connector, and therefore evenwhen only the digital ASIC malfunctions, only the digital ASIC substrateneeds to be replaced. In this way, this structure can further reducewaste.

In the above explanations, the substrate body 20 a (detector substrate20A) for mounting the detectors 21 is different from the substrate body20 b (ASIC substrate 20B) for mounting the ASICs 24, 26. Thus, when, forexample, both ASICs are soldered to a substrate by means of a BGA (BallGrid Array) using reflow, only the ASIC substrate can be soldered andthis is preferable because the semiconductor radiation detector 21 neednot be exposed to a high temperature. Of course, it is also possible toarrange all the elements 21 to 26 on the same substrate and use noconnector C1.

<<Detector Unit; Unit Construction Through Housing of CombinedSubstrate>>

Next, a unit construction by housing the above described combinedsubstrate 20 in the housing 30 will be explained. This embodimentconstructs a detector unit (twelve substrate units) 2 by housing twelvecombined substrates 20 in the housing (frame) 30. The camera 11 of thePET apparatus 1 has a structure in which 60 to 70 detector units 2 arearranged in the circumferential direction in a detachable/attachablemanner (see FIG. 12B) so as to facilitate maintenance and examination(see FIG. 2).

(Housing in Housing)

As shown in FIG. 10, the detector unit 2 is provided with a housing 30,etc., for housing or holding the above described 12 combined substrates20, a high-voltage power supply PS for supplying a charge collectingvoltage to these 12 combined substrates 20, the above described FPGA 31,signal connectors for exchanging signals with the outside and powerconnectors for receiving a power supply from the outside.

As shown in FIG. 10 and FIG. 11, the combined substrates 20 are housedin the housing 30, arranged in three rows in the depth direction(longitudinal direction of the bed 14) without overlapping with oneanother and in four rows in the width direction (circumferentialdirection of the camera 11). Namely, one housing 30 houses twelvecombined substrates 20. To realize such housing, a guide member 39consisting of four rows of guide grooves (guide rails) G1 extending inthe depth direction and arranged at appropriate intervals in thecircumferential direction is disposed in the housing 30 and fitted atthe upper end of the housing (cover) 30. The guide member 39 has anopening 40 opposed to each connector C3 of a ceiling plate 30 a in theportion of each guide groove G1. Furthermore, a bottom plate 30 b of thehousing 30 is provided with four guide members 41 having one guidegroove (guide rail) G2 extending in the depth direction arranged atappropriate intervals in the circumferential direction (see FIG. 11).The guide grooves G1, G2 have a depth corresponding to a capacity ofhousing three combined substrates 20. An end of the combined substrate20 on the ASIC substrate 20B side is housed in the guide groove G1 andan end of the combined substrate 20 on the detector substrate 20A sideis housed in the guide groove G2. Three combined substrates 20 are heldin the depth direction of the guide grooves G1, G2. Note that since theend of the combined substrate 20 on the ASIC substrate 20B side and theother end on the detector substrate 20A side are designed to be slidablein the guide grooves G1, G2, it is possible to easily position thecombined substrates 20 at predetermined locations by sliding them in theguide grooves G1, G2 with fingers, for example. In this case, eachsubstrate connector C2 is positioned in the portion of each opening 40.After a predetermined number of combined substrates 20 are arranged inthe housing 30, the ceiling plate 30 a is attached at the top end of thehousing 30 in a detachable/attachable manner using screws, etc. Eachconnector C3 fitted in the ceiling plate 30 a is inserted in thecorresponding opening 40 and connected to the corresponding substrateconnector C2. The terms “upper” and “lower” sections of the housing 30are applicable when the housing 30 is removed from the camera 11, andwhen the housing 30 is mounted in the camera 11 as shown in FIG. 2, theupper and lower sections may be inverted or turned 90 degrees to be“right” and “left” sections or located diagonally.

As shown in FIG. 11, the ceiling plate 30 a of the housing 30 isprovided with not only the four rows of guide grooves G1 but also FPGA31 and connector 38. The connector 38 is connected to the FPGA 31. TheFPGA 31 is programmable in the field. In this aspect, the FPGA 31 isdifferent from the ASIC in that it is not programmable. Therefore, aswith this embodiment, even if the number or type of the combinedsubstrates 20 to be housed changes, the FPGA 31 can be programmed in thefield to be adaptable to changes in the number of substratesappropriately.

Since the detectors 21 using CdTe as the semiconductor material S inthis embodiment generate charge in reaction with light, the housing 30is made of a material having light shielding properties such as aluminumand an alloy of aluminum and designed in such a way as to eliminate gapsthrough which light enter. That is, the housing 30 is constructed tosecure light shielding properties. If, for example, light shieldingproperties are secured by other means, the housing 30 itself need not beprovided with light shielding properties and the housing 30 can be aframe (framework) to hold the detectors 21 in a detachable/attachablemanner (e.g., no light shielding plane member (panel), etc., isrequired).

As shown in FIG. 12A, the detector unit 2 is mounted via a unit supportmember 2A. Furthermore, as shown in FIG. 12B, the detector unit 2 ismounted in the camera 11 with one end supported by the unit supportmember 2A. The unit support member 2A has a hollow disk (doughnut) shapeand is provided with many windows (as many as the detector units 2 to bemounted) in the circumferential direction of the camera 11. In order tosupport the detector units 2 at one end, a flange portion serving as astopper is provided on the front side in the axial direction of the bodyof the housing 30 of the detector unit 2. Note that the flange portionsinside in the circumferential direction become obtrusive when thedetector units 2 are arranged as dense as possible in thecircumferential direction. Therefore, it is possible to remove theobtrusive flange portions from the housing 30 and allow the flangeportions outside in the circumferential direction to remain. Or it isalso possible to provide another unit support member 2A and support bothends of the detector unit 2 by the two unit support members 2A.

In order to mount the detector units 2 in the unit support member 2A,this embodiment allows many detectors 21 to be mounted in the camera 11at a time. This can considerably shorten the time of mounting thedetectors 21 in the camera 11. Furthermore, packet data (all packet datafor all detectors 21 of a combined substrate 20) output from the datatransfer apparatus 38 of all the combined substrates 20 in the detectionunit 2 is sent from the unit combination FPGA 31 provided in thedetection unit 2 to the data processing apparatus 12. In this way, thenumber of wires through which packet data is transmitted to the dataprocessing apparatus 12 in this embodiment is also significantly reducedcompared to the case where packet data is sent from each data transferapparatus 38 of the combined substrate 20 to the data processingapparatus 12.

When the detector units 2 is mounted in the camera 11, a cover 11 a isremoved to make the unit support member 2A exposed so that the detectorunits 2 are inserted until the detector units 2 touch the flangeportions. When the detector units 2 are inserted and fitted, connectionsbetween the camera 11 and the detector units 2 are made, and signals andpower supply are connected between the camera 11 and the detector units2.

(Power Supply)

Then, the high-voltage power supply apparatus PS for supplying a chargecollection voltage will be explained. As shown in FIG. 10, the detectorunit 2 provides the high-voltage power supply apparatus PS for supplyinga charge collection voltage to each detector 21 in a space formed insidethe housing 30 on the back of the FPGA 31. This high-voltage powersupply apparatus PS receives a low voltage power supply, boosts thevoltage to 300 V using a DC-DC converter (means for boosting thevoltage, which is not shown) and supplies the voltage to each detector21. 64 detectors 21 are provided per one combined substrate 20(=detector substrate 20A) on one side, and 128 on both sides. Twelvesuch combined substrates 20 are housed in one housing 30 (that is, onedetector unit 2). Thus, the high-voltage power supply apparatus PSsupplies voltages to 128×12=1536 detectors 21.

Conventionally, a supply voltage of 300 V with extremely smallfluctuations is supplied from a precision power supply apparatus in aremote place, but since (1) when the distance from the precision powersupply apparatus increases, a wider insulating structure for highvoltage wiring is required (the insulating distance increases) and (2)the voltage fluctuates due to a temperature variation of the detectors21, there is a problem that supplying a precise voltage from theprecision power supply apparatus does not necessarily result in aprecise voltage in the part of the target detectors 21.

Furthermore, to facilitate maintenance and examination, it is alsopossible to consider providing the detector unit 2 according to thisembodiment with a power connector (not shown) and removing ahigh-voltage power line extending from the precision power supplyapparatus at this power connector. That is, according to thisembodiment, it is possible to consider supplying a high-voltage powersupply to the detector units 2 from outside the units 2 via powerconnectors. However, in the case of a high voltage of 300 V, thisresults in a problem that the size of the power connector increases inaddition to the above described problem of insulation.

According to this embodiment, the high-voltage power supply apparatus PSbuilt in the detector unit 2 is connected to an external low voltage (5to 15 V) DC power supply through the power connector 42 and connector 38provided on the ceiling plate 30 a via power wiring. A high-voltageterminal of the high-voltage power supply apparatus PS is connected totwelve connectors C3 provided on the ceiling plate 30 a through theconnector 43 provided on the ceiling plate 30 a and connected toelectrodes C of the respective detectors 21 provided on the substratebody 20 a through the connector C2 of the respective combined substrates20, power wiring (not shown) in the substrate body 20 b, connector C1and power wiring (not shown) in the substrate body 20 a. The connectorsC1, C2 include not only connectors for transmitting output signals ofthe detectors 21 but also connectors for power wiring. Since thehigh-voltage power supply apparatus PS boosts a low voltage applied fromthe power supply to 300 V using a DC-DC converter, it is possible toreduce the high-voltage section and thereby shorten the insulationdistance. That is, this eliminates the necessity for using high-voltagewiring for a portion from the connector 42 to the DC power supply. Italso facilitates maintenance, etc. For the problem with voltagefluctuations, this embodiment provides not the high-precision powersupply apparatus but the high-voltage power supply apparatus PS havingaccuracy according to a temperature fluctuation of the voltage. Thiseliminates the necessity for a high-precision power supply. Furthermore,since it is a low voltage that is received from an external powersupply, it is possible to use a small power connector to be provided forthe connector 38. Using the small power connector increases the degreeof freedom in the layout. Furthermore, since the high-voltage powersupply apparatus PS is arranged in a space formed in the housing 30 onthe back side of the FPGA 31, the arrangement of the high-voltage powersupply apparatus PS in the housing 30 makes the detector unit 2 morecompact instead of upsizing. It is also possible to directly connect thehigh-voltage power supply apparatus PS to the power wiring provided onthe substrate body 20 a through the connector, without the ceiling plate30 a. Furthermore, the power connector can also be separated from theoutput signal connector of the detector 21. This prevents noise fromentering the signal wiring from the power supply system.

Furthermore, by reducing a supply voltage to the detector unit 2, it ispossible to supply power to the high-voltage power supply apparatus PSat a low voltage through the unit combination FPGA 31 as with powersupplies to the ASICs 24, 26.

Furthermore, supplying power using the high-voltage power supplyapparatus PS eliminates the necessity for insulation from the housing(GND).

The voltage supplied from the FPGA 31 to the high-voltage power supplyapparatus PS is boosted to 300 V by a DC-DC converter (not shown) in thehigh-voltage power supply apparatus PS and after boosting, passedthrough the ceiling plate 30 a of the housing 30 and supplied from ASICsubstrate 20B→detector substrate 20A→each detector 21 for each combinedsubstrate 20. That is, the housing 30 (ceiling plate 30 a) is providedwith wiring for voltage supply (not shown) for supplying a voltage fromthe high-voltage power supply apparatus PS to each combined substrate20. Furthermore, each combined substrate 20 is provided with wiring forvoltage supply which supplies a voltage supplied from the high-voltagepower supply apparatus PS to each detector 21 via the substrateconnector C2.

Embodiment 2

A nuclear medicine diagnostic apparatus according to another embodimentwill be explained. The nuclear medicine diagnostic apparatus of thisembodiment is single photon emission computer tomography (SPECT)apparatus.

This SPECT apparatus 51 will be explained using FIGS. 13 to 15. TheSPECT apparatus 51 is provided with a pair of radiation detection blocks52, a rotary holder (body of rotation) 57, a data processing apparatus12A and a display apparatus 13. The radiation detection blocks 52 aredisposed at two positions with a 180° difference in the circumferentialdirection of the rotary holder 57. More specifically, the respectiveunit support members 56 of the radiation detection blocks 52 are mountedon the rotary holder 57 with a 180° difference in the circumferentialdirection. A plurality of detector units 2A each including twelvecombined substrates 53 are mounted on the respective unit supportmembers 56 in a detachable/attachable manner. Thus, the detectors 21 aresupported by the unit support member. The construction of each detectorunit 2A is the same as that of the detector unit 2 according toEmbodiment 1 except the construction of the combined substrate 53.

The combined substrate 53 includes a detector substrate 20A and an ASICsubstrate 53B as with the above described combined substrate 20 (FIG.14). The detectors 21 at one end of each detector substrate 20A arearranged facing the bed 14. A collimator 55 made of a radiationshielding member (e.g., lead, tungsten, etc.) is provided on eachradiation detection block 52. Each collimator 55 forms many radiationpassages through which radiation (e.g., γ-rays) passes. These radiationpassages are provided in a one-to-one correspondence with the detectors21 positioned at one end of all the detector substrates 20A of oneradiation detection block 52. All the combined substrates 53 andcollimators 55 are arranged within a light/electromagnetic shield 54mounted on the rotary holder 57. The collimator 55 is mounted in thelight/electromagnetic shield 54. The light/electromagnetic shield 54cuts off influences of electromagnetic waves other than γ-rays on thedetectors 21, etc.

When the bed 14 on which an examinee administered withradiopharmaceuticals is laid is moved, the examinee is moved between thepair of radiation detection blocks 52. When the rotary holder 57 isrotated, the detector units 2A of each radiation detection block 52revolve around the examinee. γ-rays emitted form an area in the body ofthe examinee where radiopharmaceuticals are concentrated (e.g., affectedarea) C pass through the radiation passages of the collimator 55 and areintroduced into the corresponding detectors 21. The detectors 21 outputγ-rays detection signals. These γ-ray detection signals are processed byanalog ASIC 24A and digital ASIC 26A, which will be described later.

The construction of the detector substrate 20A used in this embodiment(Embodiment 2) is the same as that in Embodiment 1 and therefore theexplanations will be omitted in this embodiment. The ASIC substrate 53Bmaking up the combined substrate 53 will be explained using FIGS. 14 and15. As with the combined substrate 20, the ASIC substrate 53B connectedto the detector substrate 20A through the connector C1 includes acapacitor 22 and a resistor 23, four analog ASICs 24A and one digitalASIC 26A for each detector 21.

One analog ASIC 24A is provided with 32 sets of analog signal processingcircuits (analog signal processing apparatuses) 33A having a slow systemand fast system. One analog signal processing circuit 33A is providedfor each detector 21. Here, the fast system includes a trigger outputcircuit 24 f which outputs a trigger signal for specifying detection ofγ-rays. As with the analog ASIC 24, the slow system is provided with acharge amplifier 24 b, a polarity amplifier 24 c, a band pass filter 24d and a peak hold circuit 24 e connected in this order. One analog ASIC24A integrates 32 sets of analog signal processing circuits 33A into oneLSI. A γ-ray detection signal which is output from the detector 21 andhas passed through the capacitor 22 and resistor 23 are guided throughthe charge amplifier 24 b, polarity amplifier 24 c and band pass filter24 d and input to the peak hold circuit 24 e. The peak hold circuit 24 eholds a peak value of the γ-ray detection signal. The γ-ray detectionsignal output from the band pass filter 24 d is input to the triggeroutput circuit 24 f. The trigger output circuit 24 f outputs a triggersignal when a γ-ray detection signal at a set level or higher is inputto remove influences of noise.

The digital ASIC 26A includes a packet data generation apparatus 34A anda data transfer circuit 37 and integrates them into one LSI. The abovedescribed trigger signal is input to the ADC control circuit 36A of thepacket data generation apparatus 34A. All the digital ASICs 26A providedon the SPECT apparatus 51 receive a 64 MHz clock signal from a clockgeneration apparatus (crystal oscillator) (not shown) and operatesynchronously. The clock signal input to each digital ASIC 26A is inputto the respective ADC control circuits 36A in all the packet datageneration apparatuses 34A. The ADC control circuit 36A identifies thedetector ID when the trigger signal is input. That is, the ADC controlcircuit 36A stores a detector ID for each trigger output circuit 24 fconnected to the ADC control circuit 36A and can identify, when atrigger signal is input from a certain trigger output circuit 24 f, thedetector ID corresponding to the trigger output circuit 24 f. The ADCcontrol circuit 36A outputs an ADC control signal including the detectorID information to the ADC 25. The ADC 25 converts the peak valueinformation output from the peak hold circuit 24 e of the analog signalprocessing circuit 33A corresponding to the detector ID to a digitalsignal and outputs it. This peak value information is input to the ADCcontrol circuit 36. The ADC control circuit 36A adds the peak valueinformation to the detector ID to generate packet data. The packet data(including detector ID and peak value information) which is the digitalinformation output from the ADC control circuit 36A of each packet datageneration apparatus 34A is input to the data transfer circuit 37. Thedata transfer circuit 37 sends the packet data output from each ADCcontrol circuit 36A to the unit combination FPGA 31 of the detector unit2A periodically. The unit combination FPGA 31 outputs the digitalinformation to the information transmission wiring connected to theconnector 38.

Packet data output from the unit combination FPGA 31 is sent to the dataprocessing apparatus 12A. A rotation angle detected by an angle gauge(not shown) connected to the rotation shaft of a motor (not shown) forrotating the rotary holder 57 is input to the data processing apparatus12A. This rotation angle indicates the rotation angle of each radiationdetection block 52 and more specifically indicates the rotation angle ofeach detector 21. Based on this rotation angle, the data processingapparatus 12A determines the position (position coordinates) of eachrevolving detector 21 on the revolving orbit. In this way, the position(position coordinates) of the detector 21 when γ-rays are detected iscalculated. Based on the calculated position of the detector 21, thedata processing apparatus 12A counts γ-rays whose peak value informationreaches and exceeds a set value. This counting is performed on each areaobtained by dividing the revolving circle into 0.5° portions relative tothe rotational center of the rotary holder 57. The peak valueinformation is an accumulated value of peak values of respective γ-raydetection signals of a plurality of detectors 21 (four detectors 21arranged on a straight line in FIG. 7A) positioned on an extension ofthe radiation passage of the collimator 55. Using the positioninformation of the detectors 21 and count value (count information) ofγ-rays when γ-rays are detected, the data processing apparatus 12Acreates tomographic information on a position at whichradiopharmaceuticals are concentrated, that is, position of malignanttumor of the examinee. This tomographic information is displayed on thedisplay apparatus 13. Information such as the above described packetinformation, count value obtained by simultaneous measurement, positioninformation of the detector 21 and tomographic information are stored inthe storage apparatus of the data processing apparatus 12.

The foregoing embodiments have described the PET apparatus 1 and SPECTapparatus 51, but the present invention is also applicable to a γcamera. Functional images obtained from the γ camera are two-dimensionaland the γ camera is provided with a collimator for regulating angles ofincidence of γ-rays. Moreover, it is also possible to adopt aconstruction of a nuclear medicine diagnostic apparatus combining thePET apparatus 1 and SPECT apparatus 51, and an X-ray CT.

Mounting (housing) of the detector unit 2 in the camera 11 is notlimited to the mounting using the above described unit support member2A, but any mounting/housing means or method can be used.

It should be further understood by those skilled in the art thatalthough the foregoing description has been made on embodiments of theinvention, the invention is not limited thereto and various changes andmodifications may be made without departing from the spirit of theinvention and the scope of the appended claims.

1-55. (canceled)
 56. A nuclear medicine diagnostic apparatus,comprising: a support member; and a plurality of detector unitsdetachably mounted on said support member, said detector units includinga plurality of semiconductor radiation detectors to which radiation isinput, an integrated circuit for processing radiation detection signalsoutput from each of said semiconductor radiation detectors, a boosterfor boosting voltage and wirings for supplying voltage from said boosterto said semiconductor radiation detectors.
 57. The nuclear medicinediagnostic apparatus according to claim 56, further comprising a bed forlaying an object to be examined, wherein said plurality of detectorunits are arranged around said bed, and said semiconductor radiationdetectors are placed near said bed and said integrated circuit is placedfar from said bed in said unit substrates.
 58. The nuclear medicinediagnostic apparatus according to claim 56, further comprising atomographic information creation apparatus for creating tomographicinformation using second information obtained based on first informationoutput from said integrated circuit.
 59. A nuclear medicine diagnosticapparatus, comprising: a support member; and a plurality of detectorunits detachably mounted on said support member, said detector unitsincluding a plurality of semiconductor radiation detectors to whichradiation is input, an integrated circuit for processing radiationdetection signals output from each of said semiconductor radiationdetectors, a booster for boosting voltage and wirings for supplyingvoltage from said booster to said semiconductor radiation detectors; acounting apparatus to which outputs from said integrated circuit areinput for counting in pairs said outputs corresponding to a pair of saidradiation detection signals detected within a set time; and atomographic information creation apparatus for creating tomographicinformation using count information output from said counting apparatus.60. The nuclear medicine diagnostic apparatus according to claim 56,further comprising a plurality of unit substrates including a pluralityof said semiconductor radiation detectors and said integrated circuitare detachably mounted on said detection units.
 61. The nuclear medicinediagnostic apparatus according to claim 56, wherein said detector unitsincludes a housing member, a plurality of unit substrates detachablyhoused in said housing member and a booster placed in said housingmember, wherein said unit substrates comprise a plurality of saidsemiconductor radiation detectors and said integrated circuit.
 62. Thenuclear medicine diagnostic apparatus according to claim 56, whereinsaid integrated circuit comprises: an analog integrated circuit forprocessing signals output from said semiconductor radiation detectors;an AD converter for converting analog signals output from said analogintegrated circuit to digital signals; and a digital integrated circuitto which signals from said AD converter are input.
 63. The nuclearmedicine diagnostic apparatus according to claim 62, wherein said analogintegrated circuit has function to amplify a signal, and said digitalintegrated circuit has function to determine a radiation detection time.64. The nuclear medicine diagnostic apparatus according to claim 60,wherein said unit substrates comprise a first substrate including saidsemiconductor radiation detectors and a second substrate including saidintegrated circuit.
 65. The nuclear medicine diagnostic apparatusaccording to claim 64, wherein said first and second substrates aredetachably connected with each other.
 66. The nuclear medicinediagnostic apparatus according to claim 65, wherein said first andsecond substrates are connected so as to overlap to each other at endportions thereof.
 67. The nuclear medicine diagnostic apparatusaccording to claim 60, wherein said semiconductor radiation detectorsare placed on both planes of said unit substrates.
 68. The nuclearmedicine diagnostic apparatus according to claim 64, wherein saidsemiconductor radiation detectors are placed on both planes of saidfirst substrate.
 69. The nuclear medicine diagnostic apparatus accordingto claim 56, further comprising: a rotation body; and a bed for layingan object to be examined, wherein said support member is placed on saidrotation body.